Typically, in CT imaging systems, a rotatable gantry includes an x-ray tube, a detector, a data acquisition system (DAS), and other components that rotate about a patient that is positioned at the approximate rotational center of the gantry. X-rays emit from the x-ray tube, are attenuated by the patient, and are received at the detector. The detector typically includes a photodiode-scintillator array of pixelated elements that convert the attenuated x-rays into photons within the scintillator, and then to electrical signals within the photodiode. The electrical signals are digitized and then received within the DAS, processed. The processed signals are transmitted via a slipring (from the rotational side to the stationary side) to a computer or data processor for image reconstruction, where an image is formed.
The gantry typically includes a pre-patient collimator that defines or shapes the x-ray beam emitted from the x-ray tube. X-rays passing through the patient can cause x-ray scatter to occur, which can cause image artifacts. Thus, x-ray detectors typically include an anti-scatter grid (ASG) for collimating x-rays received at the detector.
Imaging data may be obtained using x-rays that are generated at a single polychromatic energy spectrum. However, some systems may obtain multi-energy images that provide additional information for generating images, using dual energy or fast KV switching.
Third generation multi-slices CT scanners are typically built with detectors made of scintillator/photodiodes arrays. The detectors are positioned along an arc where a focal spot is the center of the corresponding circle. The material used in these detectors generally use scintillation crystal/photodiode arrays, where the scintillation crystal absorbs X rays and converts the absorbed energy into visible light. A photodiode is used to convert the light to an electric current. The reading is proportional and linear to the total energy absorbed.
In recent years the development of volumetric or cone-beam CT technology has led to an increase in the number of slices used in CT detectors. The detector technology used in large coverage CT enables increased coverage in patient scanning, by increasing the area exposed. In CT detectors, the increase of the number of slices results in an increase in the width of the detector in Z-axis (e.g., along a length of the patient).
The x-ray detectors of current state of the art CT systems are generally composed of a two-dimensional (2D) array of scintillating pixels, coupled to a 2D array of Si photodiodes. A typical detector includes, as examples, an array of 16, 32 or 64. However, in recent years, the need for cardiac imaging has become of increasing interest, and to include imaging of the heart within one rotation. To image the heart in one rotation, the corresponding detector size needs is approximately 140 mm to 160 mm at iso-center to cover the full organ in one rotation (equivalent to a detector with 256 slices, in this example).
However, building such a detector as single mono-structure includes significant challenges. For a detector of 256 slices and perhaps an arc of 1000 channels, a total of 256,000 pixels results in the total detector. Such a massive mono-structure inevitably will include manufacturing flaws that need repair prior to shipping as new product, such as when bad pixels are detected, or when poor image quality is observed. In addition, the detector pixels include collimating elements that are directed toward the focal spot. However, collimation may need to account for the arc in not only the channel direction, but in the Z or slice direction, as well.
Some proposed designs include using mini-modules that may be fabricated, and placed on the arc both in the channel and Z or slice directions. However, placement of mini-modules on such arcs can result in geometric challenges that can result in gaps that occur between modules.
Thus, there is a need to improve CT detectors for large Z-coverage systems.